The surgical excision of diseased tissue within the body, such as a tumor or abscess, is often complicated by the inability of the surgeon to visually differentiate the diseased tissue from the normal tissue. This problem is particularly acute in the field of surgical oncology, where small numbers of tumor cells can infiltrate areas of normal tissue both adjacent to and remote from the main tumor mass. Importantly, the failure to remove all of the diseased tissue during the procedure often results in a continuation or recurrence of the original problem.
One potential solution to this problem involves the detection of radiolabelled monoclonal antibodies and other radiopharmaceuticals, which are preferentially accumulated in diseased tissues such as cancer cells. Although intraoperative probes have been developed for use with several types of radioactive materials, the historical emphasis has been on the detection of gamma radiation in particular (gamma rays or photons). See Harris et al., Nucleonics 14:102-8 (1956); Morris et al., Phys. Med. Biol. 16:397-404 (1971); Woolfenden et al., Chest 85:84-88 (1984). Unfortunately, the prior art devices designed for use with gamma-emitting radiopharmaceuticals suffer from two significant problems: 1) the tumor-to-background ratios are non-optimal for the reliable differentiation of tumors, and 2) the detection of distant sources of gamma rays further reduces the already low tumor-to-background contrast. The longer path length of gamma radiation in body tissues creates significant background contamination from distant accumulations of the radiopharmaceutical, making the detection of nearby tagged tissues difficult or impossible.
As a result, there has been a renewed interest in developing intraoperative probes which focus primarily on the detection of beta emissions (positrons and/or electrons), particularly in light of the recent discovery of positron emitters with high affinity for cancers, such as .sup.18 F-labeled-Fluoro-2-Deoxy-D-Glucose (FDG). See Wahl et al., Cancer 67:1550-54 (1991). However, recent attempts to design an accurate beta-sensitive intraoperative probe have been complicated by the fact that positron-emitting radiopharmaceuticals such as .sup.18 F-FDG create two 511 keV annihilation photons when the positron subsequently collides with an electron. The detection of these highly penetrating gamma rays greatly reduces the observed tumor-to-background contrast gained by the use of these radiopharmaceuticals.
In response, several attempts have been made to design detectors to maximize the detected positron-to-photon ratio. One possible approach to limit the effect of the annihilation photon emissions relies upon energy discrimination to reduce the photon contribution to the signal. See Raylman et al., J. Nucl. Med. 36:1869-74 (1995). The dominant mode of interaction for the 511 keV photons produced during annihilation of the positron and electron is a Compton scattering of electrons, generally below 340 keV. In contrast, the positron interacts with the detector by producing a spectrum of energies, some of which are above the Compton edge of approximately 340 keV for annihilation photons. Accordingly, by selectively counting only those events with energies above the Compton edge, the probe becomes selectively sensitized to the electrons and positrons emitted by radiopharmaceuticals such as .sup.18 F, which create annihilation photons.
Unfortunately, most of the detectors proposed thus far for use as positron probes have utilized plastic scintillators. See Lerch et al., Am. J. Physiol. 242:H62-H67 (1982); Raylman et al., J. Nucl. Med. 35:909-13 (1994); Daghighian et al., Med. Phys. 36:1869-74 (1995). The application of the energy discrimination technique with plastic scintillators is problematic due to the poor energy resolution of this material, which is a measure of how well the energy of a specific type of radiation (such as gamma rays) can be defined. Moreover, inefficiencies in the collection of the scintillation light produced by the plastic scintillators also reduce the energy resolution of these detection devices.
An alternative method and device proposed and patented by Daghighian et al. involves the use of two separate plastic scintillation detectors, whereby the signals from the shielded outer detector are used to correct for photon contamination of the signal from the inner detector. See Daghighian et al., Med. Phys. 21:153-7 (1994); U.S. Pat. No. 5,008,546 to Mazziotta et al. Correction of signal contamination is accomplished by a weighted subtraction of the outer detector count rate from the inner detector count rate. The weighting factor is the ratio of the gamma counting efficiencies of the two detectors, which is calculated during a relatively simple calibration procedure.
While the use of a second detector to measure the background contamination is somewhat effective, this addition unfortunately results in a probe tip which is always physically larger than a single detector. Therefore, the practical application of this type of probe is problematic where space is a premium, such as with intraluminal probes and other situations where the surgical field is small. Moreover, the reduction of the surgical field continues to increase as minimally invasive surgical procedures are developed, and therefore a useful alternative to the two-detector method is needed. Furthermore, it is not clear that the background subtraction/weighting function remains constant when the probe is presented with gamma rays entering the detector volume other than through the front window. This problem is very often present in many common surgical applications.
Accordingly, there is still a substantial need in the art for an intraoperative probe which can differentiate diseased tissue based on beta emissions from a radiopharmaceutical. The probe must also have a minimal size for less intrusive operation during surgery, while at the same time provide increased sensitivity and selectivity.